Percutaneous thermal destruction (ablation) of problem myocardial tissue (arrhythmogenic focus) is a therapeutic procedure used with increasing frequency for treatment of cardiac arrhythmias (e.g., ventricular tachycardia).
Medically, ablation is covered in Ablation in Cardiac Arrhythmias, G. Fontaine & M. M. Scheinman (Eds.), Futura Publishing Company, New York, 1987. A recent review of the ablation field is given in a chapter by D. Newman, G. T. Evans, Jr., and M. M. Scheinman entitled "Catheter Ablation of Cardiac Arrhythmias" in the 1989 issue of Current Problems in Cardiology, Year Book Medical Publishers.
Catheter ablation of ventricular tachycardia was first described in 1983 as a nonsurgical method of destroying an arrhythmogenic focus. Typically, a pacing catheter is introduced percutaneously and advanced under fluoroscopic guidance into the left heart ventricle. It is manipulated until the site of earliest activation during ventricular tachycardia is found, indicating the location of problem tissue. One or more high voltage direct-current pulses from a standard defibrillator are then applied between the distal electrode of the catheter and an external large-diameter chest wall electrode. This procedure works by destroying cardiac tissue responsible for the arrhythmia.
Although this treatment is effective in some patients, there are serious drawbacks to high voltage direct-current pulses as an ablative energy source. The shock is painful, so general anesthesia is required. More importantly, the discharge produces arcing and explosive gas formation at the catheter tip. The resultant shock wave is responsible for serious side effects. The scar created by a direct-current pulse tends to have a large border zone of injured but still viable tissue that may serve as a new focus for ventricular tachycardia.
These problems have prompted a search for alternatives to direct-current pulse as a source of ablative energy. Radiofrequency (RF) energy is a promising method being investigated. (RF without qualifiers refers here to the electromagnetic spectrum from 10 kHz to 100 GHz, as per ANSI/IEEE Standard 100-1988.) Laser energy is also being considered for catheter ablation of arrhythmias (see Cohen, U.S. Pat. No. 4,685,815) but is not pertinent to the RF implementation considered here.
RF ablation using electrosurgical power units is in clinical investigation, as a safer ablation alternative to high voltage direct current pulses. At present, continuous, unmodulated current in the range of 0.5 MHz to 1.2 MHz, such as that supplied by an electrosurgical RF power supply, is applied to the endocardium via an electrode catheter in the same manner as with a direct-current pulse. Ablative injury is produced by heating, generated by an electric field emanating from the catheter electrode. There is no gas or shock wave formation, and therefore no risk of serious barotraumatic side effects. However, as discussed in more detail later, the small size of the resulting lesion remains a problem even with RF ablation.
In order to discuss and evaluate the technical state of the art of RF ablation catheters and to compare it with embodiments of this invention, one must first establish pertinent performance requirements. A general geometrical requirement of catheter-based applicators is that they must be confined in a slender cylindrical structure with a radius commensurate with the catheter diameter. Subcutaneous insertion into the heart dictates that the catheter body must be a flexible tube no more than 2 mm in diameter and about 1 meter long. The diameter is constrained by the size of blood vessels used for catheter insertion into the heart. The length is dictated by the length of the catheter inside of the patient's body plus the length of the catheter between the patient and the external equipment.
In the discussion of catheter performance which follows, it is convenient to adopt a cylindrical coordinate system with the z-axis coincident with the catheter axis and pointed toward the distal end. The radial component is in the direction normal to the catheter z-axis, and the circumferential component has a direction around the z-axis. Radius is measured from the catheter axis.
A simple cylindrical wire heat applicator antenna is shown in FIG. 1A. Applicator antenna 10 is a conductor immersed in a lossy dielectric medium which has electrical properties typical of muscle tissue. The radius of applicator antenna 10 is "a" and its height is "h". In spite of the simple geometry and low frequency approximation used in the description, FIG. 1 retains the salient features of a radial-field coupling of pacing catheters used as an RF antenna.
In FIG. 1A, RF potential V14 is applied in a unipolar manner between applicator antenna 10 and a remote boundary 15 which corresponds to a neutral electrode applied to the skin. The exact location of boundary 15 is not important to the shape of the radial electric field E near applicator antenna 10. Electric field E16 coincides with current density vector J.sub.r =.sigma.E.sub.r in the tissue, where .sigma. is the conductivity of the tissue.
Continuity of current in the cylindrical geometry in FIG. 1A results in current density J.sub.r which decreases with the inverse of the radius r: J.sub.r =J.sub.o a/r for r&lt;h and power dissipation P=J.sub.r.sup.2 /.sigma.=(J.sub.o.sup.2 /.sigma.) (a/r).sup.2. For r&gt;h the spherical geometry is a more appropriate approximation and results in J.sub.r =J.sub.o (a/r).sup.2, and the corresponding electrical power dissipation is P=J.sub.r.sup.2 /.sigma.=(J.sub.o.sup.2 /.sigma.) (a/r).sup.4. The result is that the heating of tissue, decreases with the radius within the bounds of the second to the fourth power of a/r. This behavior of the electric field applies to a conducting medium below the microwave region. In the microwave region (f&gt;900 MHz), the radial attenuation of electric field is even faster due to the "skin depth" attenuation.
Clinical experience indicates that in order to effectively ablate ventricular tachycardia, it is desirable to thermally destroy (ablate) tissue over an area of 1-2 cm.sup.2 of the myocardium (e.g., see Moran, J. M., Kehoe, R. F., Loeb, J. M., Lictenthal, P. R., Sanders, J. H. & Michaelis, L. L. "Extended Endocardial Resection for the Treatment of Ventricular Tachycardia and Ventricular Fibrillation", Ann Thorac Surg 1982, 34:538-43). As mentioned earlier, in order to accomplish percutaneous insertion into the left ventricle, the heating applicator radius is limited to 1 mm. In order to heat a 2 cm.sup.2 area, a 2 cm long applicator can be used provided an effective heating diameter of 1 cm can be reached. To overcome present shortcomings of the RF ablation method, the size of the lesion must be increased and this requires the minimization of the radial attenuation of the electric field and the associated heat dissipation.
The destructive ablation of tissue requires a temperature of approximately 50.degree. C.; this temperature defines the outer radius R of the ablation region. It is therefore desirable to heat tissue to 50.degree. C. up to 5 mm from the catheter axis. Yet at 100.degree. C., undesirable charring and desiccation takes place. So, ideally, the maximum temperature at the applicator electrode boundary should be under 100.degree. C.
Ignoring for a moment heat conduction in the tissue, the rise in tissue temperature is proportional to the electric power dissipation which in turn is proportional to the square of the current density. In order to maintain a 100.degree. C./50.degree. C. or a factor of 2 temperature ratio between the temperature at a radius of 1 mm and the temperature at a radius of 5 mm, the ratio of the power dissipation ratio should be 2 at these two distances. Yet the performance of the current density in FIG. 1A gives at best a power dissipation at the catheter surface of (R/a).sup.2 or 25, and at worse (R/a).sup.4 or 625 times more intense than heat dissipation at a 5 mm radius.
In order to examine the effect of this wide range of heat dissipation, it is useful to divide the lossy medium in FIG. 1A into three cylindrical shells: first shell R11 adjacent to the applicator antenna 10, followed by shell R12, and R13 beginning at the 10 mm radius. Since the shells are traversed by the same current, and the potential drop across the shells is additive, power delivery can be schematically represented by three resistances R11, R12, and R13, as shown in FIG. 1B, connected in series with the source of RF potential V14.
The heat required to obtain adequate ablation at a 5 mm radius tends to desiccate blood or tissue close to the applicator antenna 10, increasing the resistivity of R11. This in turn further increases the relative power dissipation in R11 in comparison with R12 and R13, until effective impedance of the desiccated region R11 becomes, in effect, an open circuit shutting off the flow of RF power to the tissue beyond R11.
This indeed is the problem with state-of-the-art RF ablation catheters which severely limits the effective heat delivery to more distant tissue. The currently used RF ablation technique, based on a surgical RF power supply and a pacing catheter, suffers from a steep temperature gradient, reported to decay as (a/r).sup.4 and has the associated problem of charring which disrupts and limits heating and ablation.
Insulation of the applicator antenna 10 from the tissue does not reduce the heat dissipation gradient: If the applicator antenna 10 is insulated from the lossy medium by a thin dielectric tube, the effect of the dielectric can be represented by capacitor (not shown) in series with the source of RF potential V14. Now the applicator must be operated at a frequency high enough so that the impedance of the sum of resistances R11 and R12 and R13 must be higher than the capacitive impedance of the dielectric tube. R11 still dominates the heat distribution.
Effective ablation heating also requires that the heating along the heat applicator axis should be as uniform as possible. Heating should then rapidly attenuate to a negligible value along the portion of the catheter acting as a transmission line.
A key improvement requirement is therefore the ability to ablate areas significantly wider than the catheter diameter, confined only to the region of the heat applicator. Heating should not be limited by charring and desiccation at the catheter boundary.
Therefore, there is a need for a catheter-compatible RF energy delivery system which dissipates heat more uniformly in the radial direction and is well defined in the z direction, thereby leading to a more accurately controlled and larger ablated region. It is also desirable to eliminate the effect of desiccation of tissue, adjacent to the electrode, on heat dissipation to surrounding tissue.
An effective cardiac ablation catheter must satisfy three additional performance requirements:
(1) The body of the catheter should act as an efficient and reproducible RF power transmission line with the heat applicator transforming the impedance of tissue (electrically a lossy medium) to match the characteristic impedance of the transmission line. PA1 (2) The detection of an endocardial potential, needed for mapping of location of the arrhythmogenic tissue to be ablated, must coexist, without interference, with the heating function. PA1 (3) All of the above must be accomplished in a flexible catheter, about 2 mm in diameter so as to allow percutaneous insertion into the left ventricle. PA1 a helix: PA1 a helix and a gap: (Stauffer et al, U.S. Pat. No. 4,825,880, Feb. 5, 1989); PA1 linear dipoles: PA1 folded dipoles: PA1 co-linear arrays: PA1 (a) Ablation heating must create significantly higher temperature differentials (30.degree. C. for ablation vs. 5.degree. C. for hyperthermia) and must operate in the presence of rapid blood flow, and therefore requires significantly higher power levels. The capabilities of the power supply and the power carrying capability of transmission lines must therefore be higher. PA1 (b) The problem of charring and desiccation, described earlier, is absent in hyperthermia, but it can be a very important obstacle in ablation. PA1 (c) Power leakage on the outside of the catheter transmission line is unimportant in hyperthermia, yet it is unacceptable in cardiac ablation. PA1 (d) Typically in heating of a tumor, an array of antennas is used and so the interaction of the antennas is important. In ablation, only a single element is used so interactive properties are unimportant.
U.S. Pat. No. 4,641,649 issued Feb. 10, 1987 to P. Walinski, A. Rosen, and A. Greenspon describes a cardiac ablation catheter consisting of a miniature coaxial line terminating in a short protruding inner conductor applicator. This system operates at 925 MHz. To applicant's knowledge, no heat dissipation profiles for the Walinski catheter are published. However, the small area of the stub-like applicator results in an E-field attenuation which is even more precipitous than in the case of the pacing catheter electrode discussed in conjunction with FIG. 1A.
Microwave ablation catheter experiments have been reported by K. J. Beckman, & J. C. Lin et al, "Production of Reversible Atrio-Ventricular Block by Microwave Energy" abstracted in Circulation 76 (IV): IV-405, 1987. Technical details of a folded dipole applicator catheter used by Beckman have been described by J. C. Lin and Yu-jin Wang in "An Implantable Microwave Antenna for Interstitial Hyperthermia" in Proceedings of the IEEE, Vol. 75 (8), p. 1132, August, 1987. The heating profile indicates an unacceptably high heat dissipation along the transmission line. Neither of the two Lin references address the all important issue of integration of monitoring of endocardial potential with the folded dipole heat applicator.
There is a large body of technical knowledge concerned with the RF catheter heating developed for oncological applications. The catheters are inserted typically to the depth of a few centimeters into a cancerous tumor and heat the tumor tissue by a few degrees centigrade. It was found that heated tumor tissue is more susceptible to chemotherapy.
A variety of oncological applicators have been proposed including:
(LeVeen, U.S. Pat. Nos. 4,154,246, 22,4,1986; PA2 Pchelnikof SU 1,266,548-A-1, Oct. 30, 1986; and PA2 Hines et al, U.S. Pat. No. 4,583,556, Apr. 22, 1986); PA2 (B. E. Lyons, R. H. Britt, and J. W. Strohbehn in "Localized Hyperthermia in the Treatment of Malignant Brain Tumors Using an Interstitial Microwave Antenna Array:, IEEE Trans on Biomedical Engineering Vol. BME-31 (1), pp. 53-62, January, 1984; PA2 (J. C. Lin and Yu-jin Wang "An Implantable Microwave Antenna for Interstitial Hyperthermia" in Proceedings of the IEEE, Vol. 75 (8): 1132, August, 1987); and PA2 (Kasevich et al, U.S. Pat. Nos. 4,700,716, Oct. 20, 1987).
RF cardiac ablation and oncological applications have the common objective of uniform heating of tissue. There are, however, a number of differences in the requirements for ablation vs. hyperthermia.
Ablation applications require uniform heating, combined with accurate monitoring of the endocardial potential, without interference and preferably without introduction of any additional catheter wires. None of the oncological references quoted above address the issue of monitoring of endocardial potential. Other differences between hyperthermia and ablation are: